This invention relates generally to methods and apparatus for generating images from data collected in a CT scan, and more particularly to generating images from data collected in a helical scan with a gantry tilted.
In at least some computed tomography (CT) imaging system configurations, an x-ray source projects a fan-shaped beam which is collimated to lie within an X-Y plane of a Cartesian coordinate system and generally referred to as the xe2x80x9cimaging planexe2x80x9d. The x-ray beam passes through the object being imaged, such as a patient. The beam, after being attenuated by the object, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is dependent upon the attenuation of the x-ray beam by the object. Each detector element of the array produces a separate electrical signal that is a measurement of the beam attenuation at the detector location. The attenuation measurements from all the detectors are acquired separately to produce a transmission profile.
In known third generation CT systems, the x-ray source and the detector array are rotated with a gantry within the imaging plane and around the object to be imaged so that the angle at which the x-ray beam intersects the object constantly changes. X-ray sources typically include x-ray tubes, which emit the x-ray beam at a focal spot. X-ray detectors typically include a collimator for collimating x-ray beams received at the detector, a scintillator adjacent the collimator, and photodetectors adjacent the scintillator.
In some clinical applications, a helical CT scan is performed with the gantry tilted. For example, when performing head scans, the CT gantry is tilted to avoid radiation to the eyes. For multi-slice helical CT, however, image artifacts will result if the projection data is not properly adjusted. For example, in multislice CT, the iso-centers defined by the gantry do not align with the image reconstruction iso-center. To avoid image artifacts, the projection data is shifted so that the gantry iso-centers coincides with the reconstruction iso-center. The amount of shift depends on the tilt angle, detector row, and the projection angle. The adjustment process is complex and requires significant computation.
To reduce the computational load in comparison to the iso-center shift described above, instead of selecting the z-axis of the coordinate system as the iso-center for reconstruction, the gantry iso-center of one of the detector rows (e.g., one of the center rows) is selected as the reconstruction iso-center. In this arrangement, the detector row for which the iso-center is based need not to undergo iso-center shift, since the defined reconstruction iso-center is the row iso-center. Data from the other detector rows is shifted relative to the selected detector row.
When the gantry iso-center of one of the detector rows (e.g., one of the center rows) is selected as the reconstruction iso-center, the reconstructed image will be shifted along yxe2x80x2-axis relative to the original compensation schemes described above. To ensure the images generated with both schemes are identical in location, an adjustment can be added in the backprojection process.
By selecting the gantry iso-center of one of the detector rows (e.g., one of the center rows) as the reconstruction iso-center as described above, and in a four row detector, only data collected from 3 out of the 4 detector rows need be shifted, which represents a 25% saving in terms of computation for the iso-center shift. For a twin scanner configuration (i.e., a detector with 2 rows of detector cells), a 50% saving in terms of computation can be realized.